1. Field of the Invention
This invention relates to a super-conducting magnet, in particular to a super-conducting magnet for use in a magnetic resonance imaging system of the type employed for medical diagnostic investigations of the interior of the human body or other tissue.
2. Description of the Prior Art
Magnet resonance imaging systems (hereinafter MRI systems) are employed to image intact biological systems and rely on nuclear magnetic resonance (hereinafter NMR). Like X-rays and ultra-sound procedures, NMR is a non-invasive analytical technique employed to examine a body. Unlike X-rays, however, NMR is a non-ionising, non-destructive process that can be employed continuously. In comparison to ultra-sound, the quality of projections or images produced with NMR are superior.
Basically, NMR is a process that results when nuclei with magnetic moments are subjected to a magnetic field. The NMR technique detects radio-frequency signals emitted from nuclei when exposed to at least two externally applied fields, one of which fields is a strong magnetic field. The magnetic field is required to polarise the nuclei. In the most common use of NMR, which permits imaging or detecting of the distribution of the water which forms some 70% of the human body, the magnetic field is required to polarise the hydrogen nuclei of the water. This, in particular, allows derivation of information of the density and chemical state of the inner soft tissues of the body.
MRI systems have been built with permanent magnets and even with resistive electro-magnets. Such magnets only produce a maximum field of 500-1000 gauss (50-100 mT) with the result that the quality and strength of the magnetic resonance signal is insufficient to obtain a fast and sufficiently detailed picture of the interior of the body. Time is relevant because movement of the patient will result in blurring of the image and for other medical reasons, such as blood flow. A long imaging period also reduces the number of patients who can be examined.
To enable an MRI system to produce faster and more detailed information, i.e. an image with more contrast and better resolution, the strength of the NMR signal must be increased. This is most readily achieved by increasing the strength of the magnetic field used. A suitable field should be at least 3000 and preferably 5000 gauss (at least 300 and preferably 500 mT). In addition, the field should be extremely stable with time and uniform over the whole area in which the body or body part is located.
Super-conducting magnets satisfy all the desiderata and, accordingly, despite their expense, the majority of MRI systems have been constructed with such magnets. U.S. Pat. No. 4,863,804 discloses superconducting wires suitable for use in a superconducting magnet. In a generally employed arrangement, the super-conducting magnet is often provided in the form of a solenoid of 1 m internal diameter, 2-3 m length and a weight of several tonnes. The magnet is often provided with a shield formed, at least externally, of iron to reduce the stray magnetic field which is a potential health hazard. WO88/08200 discloses such an integral external shield for a MR magnet.
A 600 mm internal diameter magnet would be sufficient to accommodate the average human torso but, as noted above, generally, a magnet with a 1 m internal diameter is used. This has the disadvantage that the cost of the magnet is high since, roughly, the stored energy of the coil is proportional to the cube of its inner diameter and its cost may be expected to have a similar proportionality. The reasons why super-conducting magnets of known MRI systems have an internal diameter greater than that which, in principle, is required, are discussed below.
Within the bore of an MRI system super-conducting magnet, pulse coils are fitted whose function is to alter the shape of the magnetic field while the magnetic resonance signal is being created and detected so as to provide the spacial resolution needed to create the image. The coils provide short shaped pulses of magnetic field of considerable energy content. If conductive metal parts are placed close to the exterior of these coils, heavy currents are induced which distort the shape and magnitude of the field pulses and result in a deterioration of image quality. This problem is overcome by making the main super-conducting magnet larger.
Known MRI system super-conducting magnets such as those disclosed in U.S. Pat. No. 4,782,671 and GB-A-2207813 are contained within a reservoir of liquid helium at temperature of 4K and are surrounded by thermal shields which are cooled by liquid nitrogen or by a mechanical refrigerator. Liquid helium is expensive and inconvenient to handle. Great care has to he taken to minimise heat ingress to the liquid helium which would boil it away. In order to ensure that it is absolutely lead-tight, the liquid helium reservoir is normally made of continuously welded metal. The pulse coils will tend to generate eddy currents in this reservoir because of its form, which will cause heating of the metal and loss of the liquid helium. To prevent this, the bore of the helium-cooled magnet be provided with a pulse field shield of metal held at temperature of 40 or 77K and several millimeters thick which absorbs the stray field from the coils. The eddy current heating of this pulse field shield is dissipated by a mechanical refrigerator or by use of boiling liquid nitrogen. The use of the pulse field shield introduces design problems since it must he large so as not to be inductively coupled to the pulse coils which will cause it to interfere with their operation. The need for the shield to be accommodated between the pulse coils inside the magnet and a holy to be investigated by NMR dictates the use of a magnet with an internal bore of much greater diameter than the average body.